Hearing assistance device comprising an implantable part

ABSTRACT

The application relates to a hearing assistance device comprising an implantable part for electrically stimulating an auditory nerve of a user. The application further relates to a method of operating a hearing assistance device. The object of the present application is to improve the electrical stimulation of the cochlear nerve by a cochlear implant hearing assistance device. The problem is solved in that the implanted part comprises a) a current source generator; and b) an electrode array configured to be located inside one of the cochlear scala or adjacent to the auditory nerve. The hearing assistance device is configured to produce a time-varying waveform delivered by the current source generator, the time-varying waveform comprising a positively sloping positive pulse. This has the advantages of providing a smaller spatial spread of neurons being discharged by electric stimuli from a given electrode. The invention may e.g. be used in cochlear implant hearing assistance devices.

TECHNICAL FIELD

The present application relates to hearing assistance devices andmethods of stimulating the auditory system, in particular to electricalstimulation of the cochlear nerve. The disclosure relates specificallyto a hearing assistance device comprising an implantable part forelectrically stimulating an auditory nerve of the user.

The application furthermore relates to a method of operating a hearingassistance device, the hearing assistance device comprising animplantable part.

The application further relates to a data processing system comprising aprocessor and program code means for causing the processor to perform atleast some of the steps of the method.

The scheme for providing an appropriate stimulation waveform outlined inthe present disclosure may be generally applicable to electric nervestimulation. Embodiments of the disclosure may e.g. be useful inapplications such as cochlear implant hearing assistance devices.

BACKGROUND

The following account of the background relates to one of the areas ofapplication of the present application, hearing aids for electricalstimulation of the cochlear nerve, typically termed ‘cochlear implanthearing (assistance) devices’ or simply ‘cochlear implants’ (CI).

Cochlear implant hearing assistance devices have been known in manyyears in a variety of configurations, but typically comprising

a) a number of electrodes implantable in different locations of thecochlea allowing a stimulation of different frequencies of the audiblerange,b) an external part for picking up and processing sound from theenvironment, and for determining sequences of pulses for stimulation ofthe electrodes in dependence on the current input sound,c) a (typically wireless, e.g. inductive) communication link forsimultaneously transmitting information about the stimulation sequencesand for transferring energy tod) an implanted part allowing the stimulation to be generated andapplied to the relevant of said electrodes.

Such systems are e.g. described in U.S. Pat. No. 4,207,441 and in U.S.Pat. No. 4,532,930.

A cochlear implant electrically stimulates the auditory nerve of a deafpatient to produce a sound perception. The CI-devices typically havebetween 12 and 24 processing channels which encode the sound energylevel at different cochlear locations. Due to electrical current spreadwithin the scala tympani (i.e. the cochlear duct in which the electrodearray lies), the spectral resolution is reduced and therefore the speechperception performance of cochlear implant patients does not improvewhen the number of activated channels is increased higher than appr. 8to 12. This is in contrast to normally hearing listeners which benefitfrom an increasing number of processing channels in psycho-acousticstudies using speech vocoders.

The problem is well-known and different solutions have been envisionedto solve it. Most of the work has been concerned with focusing currentusing multi-polar stimulation. This current focussing technic relies ona general principle of beam-forming using multiple sources. Beam-formingis used in various technologies such as radar or microphone array. Theresults of current focussing have been disappointing for differentreasons. First, power consumption grows linearly with the number ofchannels used. Second focusing the electric field can have subtlesub-threshold effects from the side-lobe producing the focus. Because ofthese drawbacks, other works to reduce the spread of excitation areon-going. Recently another approach has been to use optical stimulationproduced by laser pulses in the infra-red range. This new stimulationtechnique may not have the same level of spreading but is stillcurrently very power consuming.

SUMMARY

The present disclosure describes a scheme for focussed neural excitationbased on the temporal shape of stimulating waveform.

The proposed solution relies on the use of specifically designed pulseshapes, which preferably interact with the dynamics of ionic channels toonly activate a limited region of neurons around the stimulation site.

Typically, prior art electrical stimulating pulses have a square shape,which implies that the current rises instantaneously (or nearinstantaneously as limited by physical properties of the device andmedia). The effect of the pulse is stronger in front of the electrode,and the strength of the stimulation decreases with distance to thestimulation electrode. One should note that from the point of view ofthe farther away neurons, only the amplitude of the square shape pulseschanges.

For pulses with different shapes such as triangular pulses, or any pulsewith a ramp, as proposed in the present disclosure, not only theamplitude of the pulse changes (with distance), so does the slope of theramp. More specifically, the further away the neuron the shallower therising slopes of the stimulation seen by the neuron.

One has to know that certain ionic channels preclude neurons fromdischarging if the rate of depolarization of the membrane is slower thana certain “slope-threshold”. This effect has been observed in manyneurons of the auditory systems (see e.g. [Bal & Oertel; 2001]). In thatcase, the effect is due to the presence of the (low voltage activated)potassium (K+) current I_(KLVA). But other ionic currents may react inspecific fashions to the temporal shape of the stimulation waveform.

The solution to the problem above mentioned is to reduce the spread ofexcitation by using the interaction from temporal shape of a stimulationwaveform and the dynamics of ionic channels. Narrowing the spatialextent of the field of excitation in that way is simple and efficientsince it does not rely on the addition of other current sources, whichwould imply extra power consumption.

Furthermore, one may bio-engineer the neuron, either genetically orpharmacologically to express or produce specific ionic channels thatenhance or reduce this effect depending on the desired outcome.

An object of the present application is to improve the electricalstimulation of the cochlear nerve by a cochlear implant hearingassistance device.

Objects of the application are achieved by the invention described inthe accompanying claims and as described in the following.

A Hearing Assistance Device:

In an aspect of the present application, an object of the application isachieved by A hearing assistance device comprising an implantable partfor electrically stimulating an auditory nerve of a user. The implantedpart comprises

-   -   a current source generator;    -   an electrode array configured to be located inside one of the        cochlear scala or adjacent to the auditory nerve, or at the        auditory brainstem;    -   the hearing assistance device being configured to produce a        time-varying waveform delivered by said current source        generator, said time-varying waveform comprising a positively        sloping positive pulse.

This has the advantages of providing a smaller spatial spread of neuronsbeing discharged by electric stimuli from a given electrode.

In the present context, the term ‘positively sloping positive pulse’ isintended to indicate that the waveform comprises a positive pulse, whichis deliberately non-square in that it comprises a segment having apositive slope (or a finite positive tangent) followed by a fallingedge. Further a positively sloping pulse is intended to include anon-linear, e.g. a piece-wise linear, section (e.g. comprising a numberof small rising and flat phases or steps), or a section exhibiting a(e.g. monotonous) continuous functional course (e.g. exponential orlogarithmic) that on a macroscopic scale provide an increase inamplitude over the width of the positively sloped pulse.

This has the advantage of enabling a modulation of either the nervedischarge probability or the nerve discharge temporal accuracy, or bothat the same time. In particular, a narrower spatial excitation withoutthe requirement of extra current sources (which require energy to“focus” or “steer” the electric field pattern) may advantageously beprovided by embodiments of the present disclosure.

In an embodiment, the time-varying waveform comprising a positivelysloping positive pulse comprises a (first) rising edge and a (second)falling edge, wherein the height of the falling edge is larger than theheight of the rising edge. In an embodiment, the time-varying waveformcomprises a rising edge followed by a substantially monotonicallyincreasing segment followed by a falling edge. In an embodiment, therising edge has a positive slope that is larger than the slope of theintermediate segment, such as at least twice or at least 5 times aslarge. In an embodiment, the time-varying waveform comprises a risingedge followed by an intermediate segment followed by a falling edge,wherein the height of the first (rising) edge is smaller than the heightof the second (falling) edge. In an embodiment, the height of the risingedge is substantially zero. In an embodiment, the positive pulse issubstantially triangular.

In an embodiment, the time-varying waveform comprises a bi-phasicsloping, symmetric waveform stimulation pulse. In an embodiment, thetime-varying waveform comprises a negatively sloping negative pulse. Inan embodiment, the negatively sloping negative pulse comprises a firstedge and a second edge, wherein the height of the first edge is smallerthan the height of the second edge.

In an embodiment, the hearing assistance device is configured todynamically adapt the time-varying waveform to the current input signal(e.g. comprising sound from the environment or audio signals directlyreceived from an audio source). In an embodiment, the hearing assistancedevice is configured to vary the slope of the positively slopingpositive pulse in dependence of the current input signal. In anembodiment, some of the electrodes of the multi-electrode array arestimulated with time-varying waveforms comprising a sloping positivepulse while other electrodes of the multi-electrode array are stimulatedwith time-varying waveforms comprising other waveforms, e.g. a squarepositive pulse. In an embodiment, one or more specific of the electrodesof the multi-electrode array is/are stimulated with time-varyingwaveforms comprising a sloping positive pulse while in first timesegments, while being stimulated with time-varying waveforms comprisingother waveforms, e.g. a square positive pulse in second time segments.In an embodiment, said first and second time segments are determined(such as dynamically determined) according to the current input signal(e.g. its character, speech, music, noise, etc.). In an embodiment, saidfirst and second time segments are determined according to the currentacoustic environment. This has the advantage of enabling differentproperties of sound to be assigned different strategies for being“transferred” to the user.

In an embodiment, the hearing assistance device is configured todynamically adapt the time-varying waveform to the current input signal,e.g. to optimize (provide a smaller) power consumption compared to usingsquare pulses.

In an embodiment, the hearing assistance device is configured to providethat the time-varying waveform stimulation pulse is modulated in widthaccording to the frequency content of a current input signal. In otherwords, the larger the energy content of a current input signal in agiven frequency range and in a given time slot, the wider thetime-varying waveform stimulation pulse. In an embodiment, the hearingassistance device is configured to provide that the time-varyingwaveform stimulation pulse is modulated in amplitude according to thefrequency content of a current input signal. In other words, the largerthe energy content of a current input signal in a given frequency rangeand in a given time slot, the larger the amplitude of the time-varyingwaveform stimulation pulse. In an embodiment, the hearing assistancedevice is configured to provide that the time-varying waveformstimulation pulse is modulated in width as well as amplitude accordingto the frequency content of a current input signal.

In an embodiment, the current source generator is configured to delivera spatio-temporal current waveform using one or more current sources inwhich the temporal pattern is adapted to evoke a pre-defined spatialexcitation pattern (i.e. a spatial pattern of neural responseprobability or spatial pattern of neural jitter). The temporal patterncan be adapted either in an on-line (during operation of the hearingassistance device) or off-line method.

In an embodiment, the hearing assistance device comprises a multielectrode array e.g. in the form of a carrier comprising a multitude ofelectrodes adapted for being located in the cochlea in proximity of anauditory nerve of the user. The carrier is preferably flexible to allowa proper positioning of the electrodes in cochlea to achieve that theelectrodes can be inserted in cochlea. Preferably, the individualelectrodes are spatially distributed along the length of the carrier toprovide a corresponding spatial distribution along the cochlear nerve incochlea when the carrier is inserted in cochlea.

When the implanted part is operationally implanted in a person, theelectrodes are preferably located fully or partially in the cochlea ofthe person in a way allowing the electric stimulation signal to beapplied to the auditory nerve and allowing a response signal to saidstimulation (potentially) comprising a response from the nerve to bemeasured. Alternatively or additionally, the electrodes may be locatedat the auditory brainstem (to thereby allow the hearings assistancedevice to pick up evoked brainstem responses, e.g. electrically evokedauditory brain stem responses (eABRs)).

In an embodiment, the hearing assistance device consists of one fullyimplanted part only.

In an embodiment, the hearing assistance device comprises at least oneexternal part and a communications link configured to allow exchange ofdata between the external and implanted parts of the device.

In an embodiment, the hearing assistance device comprises a referenceelectrode adapted for being located outside the cochlea. In anembodiment, the hearing assistance device (e.g. the control unit) isconfigured to provide that the stimulation electrode is the same as therecording electrode. In an embodiment, the hearing assistance device(e.g. the control unit) is configured to provide that the stimulationelectrode and the recording electrode are two physically differententities.

In an embodiment, the hearing assistance device comprises more than oneelectrode array. Examples of such could be 1) a binaural case, where twoarrays (one for each of the two ears within the same device) arestimulated by the same processor, or 2) a multi-array case (for ossifiedcochleas), where several (short) electrode-arrays are used for onecochlea.

In an embodiment, the hearing assistance device is adapted to provide afrequency dependent gain to compensate for a hearing loss of a user. Inan embodiment, the hearing assistance device comprises a signalprocessing unit for enhancing the input signals and providing aprocessed output signal. Various aspects of cochlear implant hearingassistance devices are described in [Clark; 2003].

In an embodiment, the hearing assistance device comprises an inputtransducer for converting an input sound to an electric input signal. Inan embodiment, the hearing assistance device comprises a directionalmicrophone system adapted to enhance a target acoustic source among amultitude of acoustic sources in the local environment of the userwearing the hearing assistance device. In an embodiment, the directionalsystem is adapted to detect (such as adaptively detect) from whichdirection a particular part of the microphone signal originates. Thiscan be achieved in various different ways as e.g. described in the priorart.

In an embodiment, the hearing assistance device comprises an antenna andtransceiver circuitry for wirelessly receiving a direct electric inputsignal from another device, e.g. a communication device or anotherhearing assistance device. In an embodiment, the hearing assistancedevice comprises a (possibly standardized) electric interface (e.g. inthe form of a connector) for receiving a wired direct electric inputsignal from another device, e.g. a communication device or anotherhearing assistance device. In an embodiment, the direct electric inputsignal represents or comprises an audio signal and/or a control signaland/or an information signal.

In an embodiment, an analogue electric signal representing an acousticsignal is converted to a digital audio signal in an analogue-to-digital(AD) conversion process, where the analogue signal is sampled with apredefined sampling frequency or rate f_(s), f_(s) being e.g. in therange from 8 kHz to 40 kHz (adapted to the particular needs of theapplication) to provide digital samples x_(n) (or x[n]) at discretepoints in time t_(n) (or n), each audio sample representing the value ofthe acoustic signal at t_(n) by a predefined number N_(s) of bits, N_(s)being e.g. in the range from 1 to 16 bits. A digital sample x has alength in time of 1/f_(s), e.g. 50 μs, for f_(s)=20 kHz. In anembodiment, a number of audio samples are arranged in a time frame. Inan embodiment, a time frame comprises 64 audio data samples. Other framelengths may be used depending on the practical application.

In an embodiment, the hearing assistance devices comprise ananalogue-to-digital (AD) converter to digitize an analogue input with apredefined sampling rate, e.g. 20 kHz. In an embodiment, the hearingassistance devices comprise a digital-to-analogue (DA) converter toconvert a digital signal to an analogue output signal, e.g. for beingpresented to a user via an output transducer.

In an embodiment, the hearing assistance device, e.g. the microphoneunit, and or the transceiver unit comprise(s) a TF-conversion unit forproviding a time-frequency representation of an input signal. In anembodiment, the time-frequency representation comprises an array or mapof corresponding complex or real values of the signal in question in aparticular time and frequency range. In an embodiment, the TF conversionunit comprises a filter bank for filtering a (time varying) input signaland providing a number of (time varying) output signals each comprisinga distinct frequency range of the input signal. In an embodiment, the TFconversion unit comprises a Fourier transformation unit for converting atime variant input signal to a (time variant) signal in the frequencydomain. In an embodiment, the frequency range considered by the hearingassistance device from a minimum frequency f_(min) to a maximumfrequency f_(max) comprises a part of the typical human audiblefrequency range from 20 Hz to 20 kHz, e.g. a part of the range from 20Hz to 12 kHz. In an embodiment, a signal of the forward and/or analysispath of the hearing assistance device is split into a number NI offrequency bands, where NI is e.g. larger than 5, such as larger than 10,such as larger than 50, such as larger than 100, such as larger than500, at least some of which are processed individually. In anembodiment, the hearing assistance device is/are adapted to process asignal of the forward and/or analysis path in a number NP of differentfrequency channels (NP≦NI). The frequency channels may be uniform ornon-uniform in width (e.g. increasing in width with frequency),overlapping or non-overlapping.

In an embodiment, the hearing assistance device comprises a number ofdetectors of the characteristics of the current input signal (e.g. oneor more of level, frequency content, modulation, reverberation, voicecontent, noise content, auto-correlation, music, etc.). In anembodiment, the hearing assistance device comprises a classificationunit for classifying a current acoustic environment and/or the currentinput signal (e.g. fully or partially based on detectors of thecharacteristics of the current input signal).

In an embodiment, the hearing assistance device comprises a leveldetector (LD) for determining the level of an input signal (e.g. on aband level and/or of the full (wide band) signal). The input level ofthe electric microphone signal picked up from the user's acousticenvironment is e.g. a classifier of the environment. In an embodiment,the level detector is adapted to classify a current acoustic environmentof the user according to a number of different (e.g. average) signallevels, e.g. as a HIGH-LEVEL or LOW-LEVEL environment.

In an embodiment, the hearing assistance device further comprises otherrelevant functionality for the application in question, e.g.compression, noise reduction, etc.

In an embodiment, the hearing assistance device comprises a cochlearimplant hearing device.

Use:

In an aspect, use of a hearing assistance device as described above, inthe ‘detailed description of embodiments’ and in the claims, is moreoverprovided.

A Method of Operating a Hearing Assistance Device:

In an aspect, A method of operating a hearing assistance device, thehearing assistance device comprising an implantable part is furthermoreprovided by the present application. The method comprises

-   -   providing an electrode array comprising one or more stimulation        electrodes configured to be located inside one of the cochlear        scala or adjacent to the auditory nerve, or at the auditory        brainstem;    -   providing stimulation current to generate electric stimulation        pulses to one or more of said stimulation electrodes;    -   using said stimulation current to provide a parameterized        time-varying waveform of said electric stimulation pulses to one        or more of said stimulation electrodes, said parameterized        time-varying waveform comprising a positively sloping positive        pulse.

It is intended that some or all of the structural features of the devicedescribed above, in the ‘detailed description of embodiments’ or in theclaims can be combined with embodiments of the method, whenappropriately substituted by a corresponding process and vice versa.Embodiments of the method have the same advantages as the correspondingdevices.

In an embodiment, the method comprises providing a model of ioniccurrents present in a nerve or neuron, from which the temporal patternof current to deliver can be computed so that a specific dischargeprobability and/or temporal accuracy can be obtained.

In an embodiment, the method comprises one or more of the followingsteps:

-   -   Computing a first passage time probability density using a model        of said parameterized time-varying waveform.    -   Computing a set of fibres, which are activated by a single pulse        of said parameterized time-varying waveform.    -   Computing interactions between sub-sequent pulses in a pulse        train of said parameterized time-varying waveforms.

This allows a method of enhancing or creating or modifying theexpression of specific ionic channels in the neurons, which should bestimulated so that the efficiency of the temporal waveform ofstimulation may be modified or modulated in a positive or negativemanner.

The method may further facilitate one or more of the following:

-   -   a stimulation strategy in which the probability of discharge is        modulated using a parameterized pulse shape specifically        designed to limit the spread of excitation;    -   a stimulation strategy in which the parameter of the pulse-shape        is modulated to produce a specific excitation pattern by acting        on either the spread of excitation or the discharge probability        or the discharge latency;    -   a fitting procedure by which a clinician/audiologist or the        patient himself (for example through the use of custom software)        estimates the extent of the spread of excitation using the        feed-back from the patient. A number of different parameterized        time-varying waveforms may be used as electric stimulation        pulses and the user's response thereto recorded. Preferably, the        user responses for a number of identical stimulation pulses        (′played′ at different times) are recorded and averaged.

In an embodiment, the method comprises the provision of a subjectivemeasure related to the use of a masking paradigm in which the patient isasked to detect the presence of a target stimulation in concurrence witha masker presented simultaneously or earlier. The method may rely on aloudness matching method where the loudness perceived is comparedbetween pulses of differing parameterized waveforms. The fitting methodmay rely on a loudness categorisation task where the subject is asked toattribute a loudness category located on a loudness scale for pulseswith parameterized pulse-shapes.

In an embodiment, the method comprises a fitting procedure, wherein anobjective measure for fitting the pulse-shape is provided, saidobjective measure being based on recording the nerve response after itsstimulation or any evoked neural response produced by the stimulation.

In an embodiment, the method comprises exposing neurons of the user togenetic or pharmacological treatment prior to or during use of thehearing assistance device to express or produce specific ionic channelsthat enhance or reduce this effect depending on the desired outcome.

In an embodiment, the method comprises the provision of an objectivemeasure for fitting the pulse-shape, which may rely on recording thenerve response after its stimulation or any evoked neural responseproduced by the stimulation.

A Fitting System:

In an aspect, a fitting system configured to estimate the extent of thespread of excitation of different parameterized time-varying waveformsas defined in the method of operating a method of operating a hearingassistance device described above is furthermore provided by the presentdisclosure. Thereby, a predefined knowledge of correspondingparameterized time-varying waveforms and spread of excitation can be(e.g. stored in a memory of the hearing assistance device, and) utilizedin the hearing assistance device to adapt the current stimuli to thecurrent input signal.

A Method Determining a Parameterized Time-Varying Waveform:

In an aspect, a method of determining a temporal pattern of astimulation waveform is provided by the present disclosure. The methodcomprises providing a model of ionic currents present in a nerve orneuron, from which the temporal pattern of current to deliver can becomputed so that a specific discharge probability and/or temporalaccuracy can be obtained; the method comprising one or more of thefollowing steps:

-   -   Computing the first passage time probability density using a        parameterized pulse shape model.    -   Computing the set of fibres, which are activated by a single        pulse of parameterized temporal shape.    -   Computing the interactions between sub-sequent pulses in a pulse        train.

A Hearing Assistance System:

In a further aspect, a hearing assistance system comprising a hearingassistance device as described above, in the ‘detailed description ofembodiments’, and in the claims, AND an auxiliary device is moreoverprovided.

In an embodiment, the system is adapted to establish a communicationlink between the hearing assistance device and the auxiliary device toprovide that information (e.g. control and status signals, possiblyaudio signals) can be exchanged or forwarded from one to the other.

In an embodiment, the auxiliary device is or comprises an audio gatewaydevice adapted for receiving a multitude of audio signals (e.g. from anentertainment device, e.g. a TV or a music player, a telephoneapparatus, e.g. a mobile telephone or a computer, e.g. a PC) and adaptedfor selecting and/or combining an appropriate one of the received audiosignals (or combination of signals) for transmission to the hearingassistance device. In an embodiment, the auxiliary device is orcomprises a remote control for controlling functionality and operationof the hearing assistance device(s). In an embodiment, the function of aremote control is implemented in a SmartPhone, the SmartPhone possiblyrunning an APP allowing to control functionality of the hearingassistance device via the SmartPhone (the hearing assistance device(s)comprising an appropriate wireless interface to the SmartPhone, e.g.based on Bluetooth or some other standardized or proprietary scheme).

In an embodiment, the auxiliary device is another hearing assistancedevice. In an embodiment, the hearing assistance system comprises twohearing assistance devices adapted to implement a binaural hearingassistance system, e.g. a binaural hearing aid system.

DEFINITIONS

In general, a “hearing assistance device” refers to a device, such ase.g. a hearing aid or a listening device, which is adapted to improve,augment and/or protect the hearing capability of a user by receivingacoustic signals from the user's surroundings, generating corresponding(electric) audio signals, possibly modifying the audio signals, andproviding the possibly modified audio signals as audibly sensed signalsto at least one of the user's ears, e.g. (as in the present disclosure)in the form of electric signals transferred directly or indirectly tothe cochlear nerve, to other sensory nerves and/or to the auditorycortex of the user.

The hearing assistance device according to the present disclosure may beconfigured to be worn in any known way (including an implanted part),e.g. as a unit arranged behind the ear with a tube leading radiatedacoustic signals into the ear canal or with a loudspeaker arranged closeto or in the ear canal, as a unit entirely or partly arranged in thepinna and/or in the ear canal, as a unit attached to a fixture implantedinto the skull bone, as an entirely or partly implanted unit, etc. Thehearing assistance device may comprise a single unit or several unitscommunicating electronically with each other.

In some hearing assistance devices, the output unit may comprise one ormore output electrodes for providing electric signals. In some hearingassistance devices, the output electrodes may be implanted in thecochlea and/or on the inside of the skull bone and may be adapted toprovide the electric signals to the hair cells of the cochlea, to one ormore auditory nerves and/or to the auditory cortex and/or to other partsof the cerebral cortex.

A “hearing assistance system” refers to a system comprising a hearingassistance device and another device in communication with the hearingassistance device. A “binaural hearing system” refers to a systemcomprising two hearing devices and being adapted to cooperativelyprovide audible signals to both of the user's ears. In a hearingassistance system or a binaural hearing assistance system, one or bothof the hearing assistance devices may comprise other output means inaddition to output electrodes in order to provide audible signals e.g.in the form of acoustic signals radiated into the user's outer ears oracoustic signals transferred as mechanical vibrations to the user'sinner ears through the bone structure of the user's head and/or throughparts of the middle ear. In such hearing assistance devices, the outputunit may comprise an output transducer, such as e.g. a loudspeaker forproviding an air-borne acoustic signal or a vibrator for providing astructure-borne or liquid-borne acoustic signal. In a binaural hearingsystem, the output electrodes may be omitted in one hearing devicecomprising such other output means.

Further objects of the application are achieved by the embodimentsdefined in the dependent claims and in the detailed description of theinvention.

As used herein, the singular forms “a,” “an,” and “the” are intended toinclude the plural forms as well (i.e. to have the meaning “at leastone”), unless expressly stated otherwise. It will be further understoodthat the terms “includes,” “comprises,” “including,” and/or“comprising,” when used in this specification, specify the presence ofstated features, integers, steps, operations, elements, and/orcomponents, but do not preclude the presence or addition of one or moreother features, integers, steps, operations, elements, components,and/or groups thereof. It will also be understood that when an elementis referred to as being “connected” or “coupled” to another element, itcan be directly connected or coupled to the other element or interveningelements may be present, unless expressly stated otherwise.

Furthermore, “connected” or “coupled” as used herein may includewirelessly connected or coupled. As used herein, the term “and/or”includes any and all combinations of one or more of the associatedlisted items. The steps of any method disclosed herein do not have to beperformed in the exact order disclosed, unless expressly statedotherwise.

BRIEF DESCRIPTION OF DRAWINGS

The disclosure will be explained more fully below in connection with apreferred embodiment and with reference to the drawings in which:

FIGS. 1A-1C shows a use case of a hearing assistance device comprisingan implanted part according to the present disclosure, FIG. 1Aschematically showing the head of a user wearing the device, FIG. 1Bschematically showing a cross section of cochlea including a multielectrode array of the device, FIG. 1C schematically showing aperspective cross-sectional view of cochlea with the multi electrodearray is mounted in scala tympani,

FIGS. 2A-2C shows various partitions of a hearing assistance deviceaccording to the present disclosure; the embodiment in FIG. 2Acomprising only an implanted part, the embodiment in FIG. 2B comprisingan implanted part and an external part with a wireless communicationlink between them, and the embodiment in FIG. 2C comprising the sameelements as the embodiment of FIG. 2B, but where the external partcomprises an antenna part for establishing the wireless link to theimplanted part and a processing part for processing an audio signal, andwhere the antenna and processing parts are connected by a wired link,

FIGS. 3A-3E schematically shows elements of a neural response to anexemplary spatial stimulation pulse excited at a (single, mono-polar)location (L_(z)) along the cochlear nerve, FIG. 3A showing the exemplary(prior art) bi-phasic square symmetric waveform stimulation pulse with atime delay between the positive and negative phases of the pulse, thepositive and negative pulses having equal height and width, FIG. 3Billustrating stimulation of neurons along the cochlear nerve due tostimulation of a single specific electrode (E_(z)) at location L_(z),FIG. 3C schematically showing exemplary waveforms of the stimulationpulses as seen by neurons located at various locations to both sides ofthe stimulated electrode (E_(z)), FIG. 3D schematically illustrating aspatial current spread caused by the stimulation pulse at the stimulatedelectrode (E_(z)), and FIG. 3E schematically illustrating acorresponding spatial spread of neuron excitation caused by thestimulation current,

FIGS. 4A-4D schematically shows three examples of mono-polar andmulti-polar stimulation schemes, FIG. 4A illustrating the (multi-array)stimulation electrode(s) spatially distributed along a cochlear nerve,FIG. 4B illustrating a mono-polar stimulation of electrode E_(z) with abi-phasic pulse, FIG. 4C illustrating a first multi-polar stimulationcomprising stimulation of electrode E_(z) with a positive pulse andneighbouring electrodes E_(z+1) and E_(z−1) with negative pulses, andFIG. 4D illustrating a second multi-polar stimulation comprisingstimulation of electrode E_(z) with bi-phasic pulse and neighbouringelectrodes E_(z+1) and E_(z−1) with corresponding bi-phasic pulses ofopposite phase,

FIGS. 5A-5F shows six different exemplary bi-phasic stimulation ‘pulse’waveforms according to the present disclosure and their modification ata neuron located spatially apart from the primary target neurons of theelectrode emitting the stimulation pulse(s), FIG. 5A showing astimulation pulse comprising a positive sloped pulse according to thepresent disclosure and an arbitrary negative pulse (or none), FIG. 5Bshowing a bi-phasic asymmetric waveform stimulation pulse comprising apositive sloped pulse according to the present disclosure and a squarenegative pulse, FIG. 5C showing a bi-phasic asymmetric waveformstimulation pulse as in FIG. 5B, but wherein the positive and negativepulses have different widths, FIG. 5D showing a bi-phasic sloped,symmetric waveform stimulation pulse according to the presentdisclosure, FIG. 5E showing a bi-phasic asymmetric waveform stimulationpulse comprising a positive positively sloped (triangular) pulseaccording to the present disclosure and a square negative pulse, andFIG. 5F showing a bi-phasic sloped (triangular), symmetric waveformstimulation pulse according to the present disclosure,

FIGS. 6A-6B schematically shows in FIG. 6A an exemplary (step-like)relationship between the slope of a positively sloped (positive) pulse(cf. e.g. FIG. 5A) arriving at a neuron of the cochlear nerve and theprobability of discharging the neuron (cf. FIGS. 7A-7C), while FIG. 6Billustrate exemplary stimulation pulse slopes having values below andabove a threshold slope SL_(TH), respectively,

FIGS. 7A-7C shows a combined illustration of the spatial range ofexcited neurons for two different stimulation pulse waveforms, a squarewaveform as shown in FIG. 3A and a positively sloped waveform accordingto the present disclosure as shown in FIG. 5D, FIG. 7A schematicallyillustrating waveforms of the positively sloped stimulation pulses asseen by neurons located at various locations to both sides of the(single, mono-polar) stimulated electrode (E_(z)), FIG. 7B schematicallyillustrating a spatial current spread caused by the stimulation pulse atthe stimulated electrode (E_(z)), and FIG. 7C schematically illustratinga corresponding spatial spread of neuron excitation caused by thestimulation current, and

FIG. 8 shows an embodiment of a hearing assistance device comprising animplanted part partitioned as schematically illustrated in FIG. 2C, and

FIGS. 9A-9B schematically illustrates two further use cases of a hearingassistance device comprising an implanted part according to the presentdisclosure, both cases showing a bilateral (or binaural) fitting firstand second hearing assistance devices, which may not (bilateral) or may(binaural) be in communication with each other, FIG. 9A showing a usecase where each hearing assistance device comprises an implanted partaccording to the present disclosure, and FIG. 9B showing a use casewhere one of the hearing assistance devices comprises an implanted partaccording to the present disclosure, and where the other comprises anoutput transducer for mechanically or acoustically (e.g. a speaker)providing stimuli interpreted by the user as sound.

The figures are schematic and simplified for clarity, and they just showdetails which are essential to the understanding of the disclosure,while other details are left out. Throughout, the same reference signsare used for identical or corresponding parts.

Further scope of applicability of the present disclosure will becomeapparent from the detailed description given hereinafter. However, itshould be understood that the detailed description and specificexamples, while indicating preferred embodiments of the disclosure, aregiven by way of illustration only. Other embodiments may become apparentto those skilled in the art from the following detailed description.

DETAILED DESCRIPTION OF EMBODIMENTS

FIGS. 1A-1C shows a use case of a hearing assistance device comprisingan implanted part according to the present disclosure.

FIG. 1A illustrates a monaural hearing assistance system comprising asingle hearing assistance device of the cochlear implant type located ata right ear (ear1) of a user (U). Other embodiments may comprise abilateral (or binaural, or hybrid solutions, e.g. comprising twoimplanted electrodes and one common processor) system wherein a hearingassistance device is located at each of the ears (ear1, ear2) of a user(U), the two hearing assistance devices being optionally incommunication with each other in that they each comprise a transceiverfor establishing a (wireless or wired) link between them allowing thetransmission and reception of information to/from the other device. Thehearing assistance device comprises an external part and an implantedpart. Likewise, a hearing assistance system according to the presentdisclosure may additionally comprise any other multi-electrode-arraystimulation, alone or combined with any other acoustic or vibrator-basedstimulation on the same ear or the other. The external part is adaptedto be located at or in an ear of the user and comprises in theembodiment of FIG. 1A a sound capture and processing part (BTE1) adaptedto be located behind an ear of the user (U), and a communication part(COM1) in operational communication with the sound capture andprocessing part (BTE1), here via a wired connection. The communicationpart (COM1) is configured to communicate with the implanted part,including to transfer information about a current electric stimulus(e.g. representative of a current sound signal picked up by the soundcapture and processing part (BTE1)) to be applied to the cochlear nerve(cochlear nerve). The cochlear nerves are connected to the auditorycentre of the brain (the Primary Auditory Cortex, denoted PAC in FIG.1A) as indicated by the bold dashed lines. The implanted part comprisesa communication and stimulation unit (SP-TU1) and a multi-electrodearray (mea1) in operational communication with each other. Thecommunication and stimulation unit (SP-TU1) is configured to exchangeinformation with the communication unit (COM1) of the external part,including to receive stimulation information, and to generatecorresponding stimulation pulses, and to apply such pulses to electrodesof the multi-electrode array (mea1).

The multi-electrode array (mea, mea1) may comprise a flexible,originally substantially linearly shaped carrier with a number ofindividually electrically accessible electrodes located along the lengthof the carrier. In an embodiment, the flexible carrier is configured toadapt to the form of cochlea when inserted. Alternatively, themulti-electrode array (mea, mea1) may (semi-rigidly) be pre-shaped tothe form of cochlea.

FIG. 1B schematically shows a cross section of cochlea including a multielectrode array of the hearing assistance device. The multi-electrodearray (mea) is in the transversal cross-sectional view of cochlea ofFIG. 1B located at an inner wall in the right side of scala tympani. Itmay, however, be located other places in the scala tympani (e.g. asindicated by electrodes (mea2, mea3, mea4) having a dotted outline).Further, the multi-electrode array may be located elsewhere in proximityof the cochlear nerve (e.g. in one of the other scala). The three scalaof cochlea, Scala tympani, Scala media and Scala vestibuli, areschematically illustrated and denoted by the same names in FIG. 1B. Thecochlear partition (Cochlear partition) hosting (a part of) the cochlearnerve (Cochlear nerve) and separating the Scala media and Scalavestibuli from the Scala tympani, is schematically indicated in FIG. 1B.The cochlear nerve comprises hair cells (Hair cell) reaching into Scalamedia.

FIG. 1C schematically shows a perspective cross-sectional view ofcochlea (Cochlea) with the (exemplary location of) multi electrode array(mea) being mounted in scala tympani (Scala tympani). The multielectrode array (mea) comprises a carrier (carrier) comprising a numberof electrodes (electrode), e.g. 8 or more, distributed along its length(cf. dashed arrow denoted L (length) and indicating a Direction ofhelicotrema, where Scala tympani and Scala vestibuli meet). Eachelectrode (electrode) is configured to provide the option of electricalstimulation of a particular part of the cochlear nerve as indicated bythe bold line denoted electrical connections in FIG. 1C. The electricalconnections are operationally connected to the stimulation unit (SP-TU1)in FIG. 1A (or similarly to unit STU-MEU-PU-CONT in FIG. 8).

For some clinical cases of profound, deafness with non-implantablecochlea (e.g., fully ossified, Mondini syndrome, etc.) or non-stimulablecochlear nerve (e.g., nerve cut following Neurofibromatosis acoustictumor surgery), the hearing assistance device may comprise electrodesplaced on the auditory brainstem, i.e. beyond the cochlear nerve andbefore the auditory cortex. The present disclosure comprises suchembodiments where the hearing assistance device is an auditory brainstemimplant.

FIGS. 2A-2C shows various partitions of a hearing assistance deviceaccording to the present disclosure.

FIG. 2A shows a hearing assistance device (HAD) in its most basic formcomprising only a, preferably self-contained (e.g. battery driven, andcomprising an input transducer, e.g. a microphone, and appropriateprocessing capability), implanted part (IMPp). FIG. 2B shows a hearingassistance device (HAD) comprising an implanted part (IMPp) and anexternal part (EXTp) with a wireless (e.g. inductive) communication link(Wireless link) between them. The external part (EXTp) may e.g. comprisean input transducer, e.g. a microphone, and a signal processing unit forenhancing a received electric input signal and possibly for preparing ascheme for stimulating electrodes of the implanted part (IMPp) independence of the current input signal. The external part (EXTp) mayfurther comprise antenna and transceiver circuitry for transferringstimulation information (and possibly corresponding energy) to theimplanted part (IMPp) (which comprises corresponding antenna andtransceiver circuitry to allow reception of the transmitted signals andenergy, to establish the Wireless link). Alternatively, the link fromthe external part (EXTp) to the implanted part (IMPp) may be based on awired connection. FIG. 2C shows a hearing assistance device (HAD) as inFIG. 2B but where the external part (EXTp) comprises an antenna part(ANTp) for establishing the wireless link to the implanted part (IMPp)and a processing part (BTEp) for processing an audio signal, and wherethe antenna and processing parts are connected by a wired link (Wiredlink, e.g. a cable). In an embodiment, the processing part (BTEp) isconfigured to be located at an ear of the user. Alternatively, theprocessing part (BTEp) and the antenna part (ANTp) may be connected by awireless link. This may be particularly relevant, if the processing part(BTEp) is located elsewhere than at an ear of the user.

FIGS. 3A-3E shows elements of a neural response to an exemplary spatialstimulation pulse excited at a (single, mono-polar) location (L_(z))along the cochlear nerve.

FIGS. 3A-3E schematically illustrates how excitation of neurons spreadsaround an electric stimulation intended to stimulate a specific location(area) of the cochlear nerve in an exemplary electrical stimulation of acochlear implant, for which a multi-electrode array (mea in FIGS. 3B-3C)produces a stimulation (of bi-phasic symmetric, square pulse shape, cf.FIG. 3A).

FIG. 3A shows an exemplary time-variant, (prior art) bi-phasic, squaresymmetric waveform stimulation pulse with a time delay (ΔT_(pn)) betweenthe positive (SQP_(a)) and negative (SQP_(c)) phases of the pulse, thepositive (Positive pulse) and negative (Negative pulse) pulses havingequal height (A_(s)) and width (p_(wa)=p_(wc)) to conserve chargeneutrality (as indicated by the areas enclosed by the respective pulsesbeing equal: Area(AP)=Area(CP)). The time variant waveforms are drawn inan amplitude (Intensity, e.g. charge density or current (e.g. in unitsof A)) versus time (T) plot.

The solid line bi-phasic, symmetric stimulation pulse of FIG. 3A(denoted SQP_(a)(L_(z)) exhibits a square waveform of a given amplitude(A_(s), intensity, e.g. charge/phase) at the location (L_(z)) of theneurons intended for receiving pulses from the electrode in question(cf. E_(z) in FIG. 3B). When a given stimulation pulse arrives atneurons located a distance (L_(z)+/−ΔL) away from the target neurons theamplitude (intensity) of the square pulse has been modified (decreasedto A_(n), cf. dashed waveforms SQP_(a)(L_(z)+/−ΔL) andSQP_(c)(L_(z)+/−ΔL) in FIG. 3A for the positive and negative phases,respectively).

FIG. 3B schematically illustrates stimulation of neurons (neurons) alongthe cochlear nerve (Cochlear nerve) due to stimulation of a singlespecific electrode (E_(z)) at location L_(z). The current provided bythe stimulation pulse(s) is indicated by ions (encircled +, − signs,respectively, in FIG. 3B). A multi-electrode array (mea) isschematically shown along the cochlear nerve with an accompanying lengthindication (arrow L) increasing towards the (rounded off) tip (tip) ofthe carrier. Electrodes E_(z−1), E_(z), E_(z+1), . . . , E_(Ne), whereN_(e) is the number of electrodes on the carrier (carrier), are spacedapart, e.g. according to a scheme adapted to a particular user, or to ageneral user. In an embodiment, the electrodes are regularly spaced by apredefined distance (Δd_(e)). A reference electrode (Ref), e.g. to pickup charges during a mono-polar stimulation is shown. The referenceelectrode is preferably located outside cochlea.

FIG. 3C schematically shows exemplary waveforms of the stimulationpulses as experienced by neurons located at various locations to bothsides of the stimulated electrode (E_(z)). The amplitude of thestimulation pulses is decreasing with increased distance from thestimulated electrode, as indicated by the graphs in dotted ellipticalenclosures below the neurons (neurons) of the cochlear nerve in FIG. 3C.Corresponding amplitudes of the pulses are denoted A⁻², A⁻¹, A₀, A₊₁,A₊₂, respectively.

FIG. 3D schematically illustrates a spatial current spread (SCSP) causedby the stimulation pulse at the stimulated electrode (E_(z) in FIG. 3B,3C). The graph in FIG. 3D shows current versus distance (L) with adecrease in current to both sides of a maximum at the location L_(z) ofthe stimulating electrode. FIG. 3E schematically illustrates acorresponding spatial spread of neuron excitation (Neural response)caused by the stimulation current. The graph in FIG. 3E shows neuralresponse (NRES) versus distance (L) with a decrease in response to bothsides of a maximum at the location L_(z) of the stimulating electrode. Athreshold value A_(TH) indicates a level below which the neurons willnot discharge. A corresponding spatial spread ΔL(A) around the locationL_(z) of the stimulating electrode is indicated. The threshold valueA_(TH) (and hence the spatial spread ΔL(A) of the neural response)depends on characteristics of the stimulation pulse, as indicated by thegraphical insert of the bi-phasic, symmetric, square waveform associatedwith the dashed line indicating the threshold value A_(TH).

In the example of FIGS. 3A-3E, a mono-polar stimulation using a singlestimulation electrode (E_(z)) and a reference electrode (Ref) isassumed. Further, a bi-phasic, symmetric square stimulation pulse isused for illustration. However, bi-polar (or multi-polar in general)stimulation and/or single phase (positive) or asymmetric stimulation mayjust as well be used (cf. FIGS. 4A-5F). According to the presentdisclosure, a positively sloped positive stimulation pulse is preferablyused (cf. FIGS. 5A-5F). As indicated in FIG. 3C, the pulses experiencedby neurons farther away from the stimulated electrode (E_(z) in FIGS.3A-3E) decrease in amplitude with distance from the stimulated electrode(cf. (A_(i), i=−2, −1, 0, +1, +2), but may still be large enough to beperceived by (i.e. to excite or fire) neurons at such locations, asillustrated in FIG. 3D, 3E.

FIGS. 4A-4D shows three examples of mono-polar and multi-polarstimulation schemes. FIG. 4A illustrates the (multi-electrode array)stimulation electrode(s) (mea) spatially distributed (L) along acochlear nerve comprising neurons (neurons) to be stimulated. Areference electrode (Ref) for use in mono-polar stimulation is furthershown. FIG. 4B illustrates a mono-polar stimulation of electrode E_(z)with a bi-phasic pulse (utilizing reference electrode Ref for the returncurrent) in an amplitude A(E_(z)) versus time t plot. FIG. 4Cillustrates a first multi-polar (asymmetric) stimulation comprisingstimulation of electrode E_(z) with a positive pulse (A(E_(z)) versustime t) and stimulating neighbouring electrodes E_(z+1) and E_(z−1) withnegative pulses (A(E_(z+1)) and A(E_(z−1)), respectively, versus timet). To conserve charge balance, the total charge of the combinednegative phases of stimulation pulses at electrodes E_(z+1) and E_(z−1)is equal to the charge of the positive stimulation pulse at electrodeE_(z) (as indicated by the corresponding (hatched) areas of the pulses).FIG. 4D illustrates a second multi-polar stimulation comprisingstimulation of electrode E_(z) with bi-phasic pulse and neighbouringelectrodes E_(z+1) and E_(z−1) with corresponding bi-phasic pulses ofopposite phase. Again, charge neutrality is intended as indicated byequality of the total areas of the positive phases and the total areasof the negative phases, respectively. The above stimulation schemes areonly examples of mono-polar and multi-polar stimulation. Any mono-polarand multi-polar stimulation scheme (providing charge neutrality) may beused in combination with the stimulation waveform according to thepresent disclosure.

From a general stimulation point of view (spatial and temporaldefinitions), the polarity will change due to both temporal and spatialdefinition. I.e., a positive temporal waveform definition can beinversed with polarity inversion of the electrode from a spatialdefinition criterion, in other words, physically inversed.

FIGS. 5A-5F shows six different exemplary bi-phasic stimulation ‘pulse’waveforms according to the present disclosure and their modification ata neuron located spatially apart from the primary target neurons of theelectrode emitting the stimulation pulse(s).

FIG. 5A shows a parameterized time-varying waveform (Intensity versusTime) comprising a positively sloping positive pulse stimulation pulseaccording to the present disclosure and an arbitrary negative pulse (ornone).

The (optional) negative pulse is shown in the dotted box denotedNegative pulse. A number of purely exemplary waveforms of the (optional)negative pulse are indicated in dashed line in the dotted box, includinga passive discharge waveform. The parameterized time-varying waveform ofthe positive pulse (Positive pulse, SLP_(a)(L_(z)), solid line waveformin FIG. 5A) comprises a positively sloped waveform defined by heights ofthe rising (A_(s1)) and falling (A_(s2)) edges of the positive pulse(resulting in slope angle α) at the location (L_(z)) of stimulation. Thepositive pulse has a width in time of p_(wa). When a given positivestimulation pulse arrives at neurons located a distance (ΔL) away fromthe target neurons the amplitudes and the slope of the sloped pulse hasbeen modified (both decreased, to (A_(n1), A_(n2)) and β, respectively,cf. dashed line waveform SLP_(a)(L_(z)+/−ΔL) in FIG. 5A). In general,the width of the positive pulse (p_(wa)) is adapted to the currentapplication (and possibly dynamically adapted to the current need forstimulation in the frequency range aimed at by a particular electrode).In an embodiment, the width of the positive pulse is of the order oftens of microseconds. In an embodiment, the width of the positive pulseis larger than 5 μs. In an embodiment, the width of the positive pulseis smaller than 100 μs. The negative pulse(s) is preferably configuredto maintain charge neutrality together with the positive pulse(s). Thenegative pulse(s) may—together with the positive pulse—form part of abiphasic pulse in a mono-polar stimulation configuration, or may beapplied to another electrode in a multi-polar stimulation configuration(cf. e.g. FIGS. 4A-4D).

The parameterized time-varying waveform stimulation pulse of FIG. 5B isbi-phasic and asymmetric in that it comprises different positive andnegative pulse waveforms. The positive stimulation pulse (Positivepulse, SLP_(a)(L_(z)), solid line waveform) exhibiting a positivelysloped waveform is equal to that of FIG. 5A. The negative stimulationpulse (Negative pulse, SQP_(C)(L_(z)), dashed line waveform) is a squarepulse. The pulse width (p_(wc)) of the negative phase is equal to thepulse width (p_(wa)) of the positive phase. Preferably, the area(charge) of the positive and negative pulses are equal (or a differenceis otherwise compensated for, e.g. by multi-polar stimulation). Thepositive and negative phases are separated by a time delay ΔT_(pn). Inan embodiment, the time delay is zero (see e.g. FIG. 5C).

FIG. 5C shows a bi-phasic asymmetric waveform stimulation pulse as inFIG. 5B, but wherein the positive and negative pulses have differentwidths (p_(wa)<p_(wc)) and where the time delay between the positive andnegative phases is minimal (e.g. intended to be zero). Again,preferably, the area (charge) of the positive and negative pulses areequal.

FIG. 5D shows a bi-phasic sloped, symmetric waveform stimulation pulseaccording to the present disclosure comprising an arbitrary time delay(ΔT_(pn)) between the positive and negative phases of the bi-phasicpulse. The positive pulse is as shown in FIGS. 5A, 5B and 5C, and thenegative pulse is a symmetrically generated version thereof (e.g.mirrored around a horizontal axis). Hence, the area (charge) of thepositive and negative pulses are equal, thereby preserving chargeneutrality.

FIG. 5E shows a bi-phasic asymmetric waveform stimulation pulsecomprising a positive positively sloped (triangular) pulseTRP_(a)(L_(z)) according to the present disclosure and a square negativepulse SQP_(c)(L_(z)) comprising an arbitrary time delay (ΔT_(pn))between the positive and negative phases of the bi-phasic pulse. Thetriangular pulse is a special case of the parameterized time-varyingwaveform stimulation pulse of FIG. 5A, where the vertical rising edge isabsent (A_(s1)=A_(n1)=0). Otherwise, it behaves as previously described,e.g. in connection with FIG. 5A. As in FIG. 5C, the positive andnegative pulses have different widths (p_(wa)<p_(wc)). Again,preferably, the area (charge) of the positive and negative pulses areequal.

FIG. 5F shows a biphasic sloped (triangular), symmetric waveformstimulation pulse according to the present disclosure comprising anarbitrary time delay (ΔT_(pn)) between the positive and negative phasesof the bi-phasic pulse. The positive pulse is a triangular pulse asshown in FIG. 5E and the negative pulse is a symmetrically generatedversion thereof. Hence, the area (charge) of the positive and negativepulses are equal, thereby preserving charge neutrality.

To summarize FIGS. 5A-5F: According to the present disclosure, animportant property of the stimulation pulse is the temporal shape of thepositive phase. A possible time lag between the positive and negativephases and the waveform of the negative phase are of minor importance.An advantage of the present, sloped stimulation pulse scheme is that itallows to use a variation of slope to code for intensity. Square pulsesallow intensity coding too. Preferably, a combination of a square pulseand a triangular pulse (here termed a ‘sloped pulse’, cf. FIG. 5B, 5C,5D) can be used. A slope could be a constantly rising current as shownin FIGS. 5A-5F, or a fast succession of flat and rising current like astair, or any other appropriate increase of the intensity from a lowerstart value to a higher end value.

The goals are:

-   -   to reduce the spatial current spreading (improve spatial        selectivity)    -   to improve intensity coding (e.g. have same amplitude coding        with less energy using sloped instead of square pulses).

For sloped pulses (cf. FIGS. 5A-5F), a neuron located a distance fromthe neuron(s) intended for stimulation sees a smaller pulse. However, inaddition to a smaller amplitude of the pulse, the slope of the pulse(stimulating current) has also decreased.

It is assumed that neurons in the auditory system are sensitive to therate of depolarization. This means that they will discharge only if theslope of stimulation is higher than a certain value. This is assumed tobe due to the presence of a fast activating sub-threshold potassiumchannel.

Because of this rate threshold, using a pulse with ramp (a sloped pulse)in its temporal profile will allow to reduce the stimulation spatialselectivity (as illustrated in FIGS. 7A-7C).

FIGS. 6A-6B shows in FIG. 6A an exemplary (step-like) relationshipbetween the slope of a positively sloped (positive) pulse (cf. e.g. FIG.5A) arriving at a neuron of the cochlear nerve and the probability ofdischarging the neuron (cf. FIGS. 7A-7C), while FIG. 6B illustrateexemplary stimulation pulse slopes having values below and above athreshold slope SLTH, respectively. It is believed that the observedproperty of neurons in the auditory system to be dependent on the slopeof the stimulation pulses is linked to the presence of the (low voltageactivated) potassium (K+) current I_(KLVA).

FIGS. 7A-7C is a combined illustration of the spatial range of excitedneurons for two different stimulation pulse waveforms, a square waveformas shown in FIG. 3A and a positively sloped waveform according to thepresent disclosure as shown in FIG. 5D. FIG. 7A (corresponding to FIG.3C dealing with the same issue but for a prior art, square waveform)schematically illustrates waveforms of the positively sloped stimulationpulses as seen by neurons located at various locations to both sides ofthe (single, mono-polar) stimulated electrode (E_(z)). As also indicatedin FIGS. 5A-5F, the amplitudes (A_(i), i=−2, −1, 0, +1, +2) AND slopes(β_(i), i=−2, −1, 0, +1, +2) of the pulses decrease with increasingdistance from the stimulation electrode (E_(z)). FIG. 7B schematicallyillustrates a spatial current spread (SCSP) caused by the stimulationpulse at the stimulated electrode (E_(z)). The graph in FIG. 7B showscurrent (SCSP) versus distance (L) with a decrease in current to bothsides of a maximum at the location L_(z) of the stimulating electrode.FIG. 7C schematically illustrates a corresponding spatial spread ofneuron excitation (Neural response) caused by the stimulation current byillustrating a probability of neuron excitation along a length of thecochlear nerve centred around a location (L_(z)) of a stimulatedelectrode (E_(z)) for two different bi-phasic simulation pulse waveforms(as illustrated in FIGS. 3A and 5D, respectively, and indicated byinserts associated with the two different neural response curves),resulting in a different spread ΔL of excitation of neurons for the twowaveforms. The threshold value A_(TH) indicating a level below which theneurons will not discharge (dashed line). Corresponding spatial spreadsΔL(SQ) and ΔL(SL) around the location L_(z) of the stimulating electrodeis indicated for each the two stimulation pulse waveforms, square (SQ)and sloped (SL), respectively. As indicated in FIG. 7C, the spatialspread ΔL(SQ) of the neural response of the square stimulation pulsewaveforms is larger than the spatial spread ΔL(SL) of the neuralresponse of the sloped stimulation pulse waveforms.

FIG. 8 shows an embodiment of a hearing assistance device comprising animplanted part partitioned as schematically illustrated in FIG. 2C. FIG.8 illustrates a ‘normal operation scenario’, where electrodes (E_(z),z=1, 2, . . . , N_(e)) of a flexible multi-electrode array (mea) of theimplanted part (IMPp) (inserted into one of the scala of cochlea, e.g.scala tympani, and having its electrodes distributed along the extent ofthe cochlear nerve). The individual electrodes (E_(z)) are stimulated independence of an acoustic input signal (AlnS) picked up by a microphoneof an external part (EXTp) of the system (cf. FIG. 2B or 2C, hereexternal part BTEp, e.g. adapted for being located behind an ear of auser). In the embodiment of FIG. 8, the relevant current stimulationscheme generated in the external BTEp part and the accompanyingnecessary electric energy are transferred to the implanted part via acommunication link (Com-Link) between the implanted part (IMPp) and anexternal antenna part (ANTp).

The external BTEp part comprises a forward signal path comprising:

-   -   a microphone (or microphone system, e.g. for providing        directionality in a specific DIR-mode),    -   an A/D converter (A/D) for converting an analogue input signal        to a digital signal by sampling the analogue input signal with a        sampling frequency f_(s),    -   a pre-emphasis filter (PEF) (e.g. a FIR filter) for adapting the        input levels to a loudness perception of a normally hearing        person (psychoacoustic adaptation),    -   an analysis filter bank (A-FB) for converting a single time        variant input signal to time-variant signals in a number p of        frequency bands (I₁:I_(p)). The analysis filter bank may e.g.        comprise a 128 point FFT providing p=64 frequency bands (or        alternatively a filter bank followed by an envelope detector),    -   a regrouping unit (REGR) for allocating p frequency bands to a        number q of channels (CH₁:CH_(q)) equal to the number of        electrodes used, e.g. q=20, configurable based on user data (cf.        unit User specific data), e.g. based on the Bark scale or        ‘critical bands’),    -   a noise reduction algorithm (NR, with settings based on User        specific data) adapted to attenuate signal components that are        judged not to be part of a target signal, the noise reduction        algorithm e.g. working independently on signals of each channel        (CH₁:CH_(q)),    -   a compression scheme (COMP, with settings based on User specific        data) adapted to provide a level dependent compression of an        input signal of each channel (CH₁:CH_(q)),    -   a stimulation generator (STG) for generating a representation of        the stimuli corresponding to a given intensity in a given        frequency range at a given point in time (reflecting the current        input audio signal) to be applied to corresponding electrodes of        the implanted part,    -   a local energy source (BAT), e.g. a battery, such as a        rechargeable battery for energizing components of the hearing        assistance device (BTEp, ANTp, IMPp), and    -   a stimulus data coding unit (COD-PLS, with settings based on        User specific data) for generating a scheme (incl. providing        energy for stimulating each of the (active) electrodes (E_(z),        max q electrodes, typically less) of the implanted part (IMPp),        and forwarding stimuli (or coded stimuli) and energy via a cable        to the antenna part (ANTp).

The unit User specific data) may represent user data stored in a memoryof the BTEp part or user data read into the various algorithms during afitting session (or a combination of the two). Such data may includefrequency dependent hearing thresholds and uncomfort levels (related toelectric stimulation of the individual electrodes). The user specificdata may include age, gender, etc.

In an alternative embodiment, the components of the external part (BTEp)are included in the implanted part (IMPp), whereby the hearingassistance device is self-contained (cf. FIG. 2A). In such anembodiment, only a communication link to an external fitting system isnecessary.

In the embodiment of FIG. 8, a cable (denoted Cable to ANTp, and Cablefrom BTEp, in the BTEp- and ANT-p-ends, respectively) connects theBTE-part (BTEp) to the antenna part (ANTp). The cable provides separatedigital data and power (denoted Stimuli-data+power) to the antenna part(ANTp).

The antenna part (ANTp) is adapted for being located at the ear of theuser allowing a communication link (Com-link) to be established with theimplanted part (IMPp). The antenna part comprises:

-   -   a power and data mixing unit (e.g. incl. a crystal oscillator)        forming part of    -   an inductive transmitter (and backlink receiver), (TX(Rx)) and        antenna coil (Ant).

The implanted part (IMPp) comprises:

-   -   an inductive antenna coil (Ant) and receiver (and backlink        transmitter), (RX(Tx)),    -   a multi-electrode array (mea) comprising a (typically flexible)        carrier (e.g. of silicone rubber) with a multitude of electrodes        (E_(z)) (of a corrosion resistant, e.g. noble, metal), each        being individually connectable to a current source of a        stimulation unit (STU) and preferably a voltage measurement unit        for capturing a nerve response by a capacitor:    -   a stimulation unit (STU) comprising        -   a data extraction circuit, for extracting configuration data            and stimuli data        -   a current generator for generating a stimulus current (based            on the extracted stimulus data) to be applied to the            electrodes (E_(z)),    -   an interface to the electrodes (E_(z)) comprising capacitors and        switches (SW) for switching between individual electrodes and        their connection to the stimulation unit (STU) and to a        measurement unit (MEU),    -   an operational amplifier (e.g. forming part of the measurement        unit MEU) and preferably a processing unit (PU) for processing        and identifying nerve response measurements (e.g. eCAPs), and    -   a control unit (CONT) configured to control the timing and        waveform of the application of stimulation signals in a        stimulation time period and the coupling (via switch unit (SW))        of a relevant stimulation electrode to the stimulation unit        (STU) and the optional measurement of a resulting response in a        measurement time period and the optional coupling (via switch        unit (SW)) of a relevant recording electrode to the measurement        unit (MEU).

An inductive, preferably bi-directional, communication link (Com-link)(e.g. comprising a 4 MHz carrier) is established by the inductive coils(Ant) of the antenna part (ANTp) and the implanted part (IMPp) when thetwo are located in an operational position (e.g. near the ear, on eachside of the skin of a person). A back-link from the implanted part tothe antenna- (and BTE-) part can e.g. be based on ‘load communication’.Due to the inductive coupling between the two antenna coils, any draw ofcurrent in the implanted part can be sensed in the antenna part. Therebydata-messages can be transmitted to the processor of the BTE-part (e.g.implant status signals (e.g. power level), electrode measurement data(impedances, and eCAPs). The backlink data can e.g. be coded in thesignal using pulse width modulation (PWM) or amplitude modulation (AM).Alternatively, a digital coding scheme can be applied

The external parts (BTEp and ANTp) can be partitioned in any otherappropriate way than shown in FIG. 8. In an embodiment, the outputs ofthe BTE part (BTEp) are a) digitally coded data representing theelectrode stimuli and b) a battery voltage, whereas the antenna part(ANTp) comprises a 4 MHz crystal oscillator whose output is mixed withthe coded data to provide an on-off-coded signal, which is transmittedto the implant receiver via the inductive link. In an embodiment, allnon-implanted parts of the hearing assistance device are located in asingle external device (EXTp) and a communication link (Wireless link)between the implanted and external parts allowing the necessary exchangeof information between the two parts (and possibly between the implanted(and/or the external) part and a fitting system), see e.g. FIG. 2B.

In a fitting situation or during operation, nerve responses (e.g. eCAPs)and/or electrode impedance measurements are communicated to a fittingsystem for setting up the hearing assistance device according to auser's particular needs, either directly via the antenna part (ANTp) orvia the BTE part (BTEp).

The analogue electric signal representing an acoustic signal from themicrophone is converted to a digital audio signal in theanalogue-to-digital converter (A/D). The analogue inputs signal issampled with a predefined sampling frequency or rate f_(s), f_(s) beinge.g. in the range from 8 kHz to 48 kHz (adapted to the particular needsof the application) to provide digital samples x_(n) (or x[n]) atdiscrete points in time t_(n) (or n), each audio sample representing thevalue of the acoustic signal at t_(n) by a predefined number N_(s) ofbits, N_(s) being e.g. in the range from 1 to 16 bits. A digital samplex has a length in time of 1/f_(s), e.g. 50 μs, for f_(s)=20 kHz. In anembodiment, a number of audio samples are arranged in a time frame. Inan embodiment, a time frame comprises 64 audio data samples. Other framelengths may be used depending on the practical application.

In an embodiment, the analysis filter bank (A-FB) comprise(s) aTF-conversion unit for providing a time-frequency representation of aninput signal. In an embodiment, the time-frequency representationcomprises an array or map of corresponding complex or real values of thesignal in question in a particular time and frequency range. In anembodiment, the TF conversion unit comprises a filter bank for filteringa (time varying) input signal and providing a number of (time varying)output signals each comprising a distinct frequency range of the inputsignal. In an embodiment, the TF conversion unit comprises a Fouriertransformation unit for converting a time variant input signal to a(time variant) signal in the frequency domain. In an embodiment, thefrequency range considered by the hearing assistance device from aminimum frequency f_(min) to a maximum frequency f_(max) comprises apart of the typical human audible frequency range from 20 Hz to 20 kHz,e.g. a part of the range from 20 Hz to 8 kHz, e.g. 400 Hz to 6 kHz.

FIGS. 9A-9B schematically illustrates two further use cases of a hearingassistance device comprising an implanted part according to the presentdisclosure, both cases showing a bilateral (or binaural) fitting firstand second hearing assistance devices (which may not (bilateral) orwhich may (binaural) be in communication with each other). FIG. 9A showsa use case where each hearing assistance device comprises an implantedpart according to the present disclosure. The functional parts of theindividual first and second hearing assistance devices are discussed inconnection with FIG. 1A and FIG. 8. FIG. 9B shows a use case (aso-called bimodal configuration) where one of the hearing assistancedevices (the second) comprises an implanted part according to thepresent disclosure, and where the other hearing assistance device (thefirst) comprises an output transducer (OT1) for mechanically oracoustically (e.g. a speaker) providing stimuli intended to beinterpreted by the user as sound. The output transducer (OT1) of thefirst hearing assistance device in FIG. 9B is shown as a (loud)speakerfor generating acoustic stimuli, but may alternatively or additionallycomprise a vibrator for mechanically exciting bones of the user (e.g.the skull). In an alternative embodiment, one of the hearing aid devicesmay comprise a speaker as well as an implanted part comprising amulti-electrode array. The first hearing assistance device may comprisea normal air conduction type hearing assistance device. The functionalparts of the second hearing assistance device (comprising an implantedpart) are discussed in connection with FIG. 1A and FIG. 8. The first andsecond hearing assistance devices may be configured to be able toexchange information between them. In an embodiment, first and secondhearing assistance devices each comprises transceiver units allowing awired or wireless link to be establish between them. An advantage ofusing a hearing assistance device according to the present disclosure ina bimodal fitting is that is that an improved frequency resolution ofthe implanted device can be provided to better match the frequencyresolution of the corresponding air conduction hearing device.

The invention is defined by the features of the independent claim(s).Preferred embodiments are defined in the dependent claims. Any referencenumerals in the claims are intended to be non-limiting for their scope.

Some preferred embodiments have been shown in the foregoing, but itshould be stressed that the invention is not limited to these, but maybe embodied in other ways within the subject-matter defined in thefollowing claims and equivalents thereof.

REFERENCES

-   U.S. Pat. No. 4,207,441-   U.S. Pat. No. 4,532,930-   [Clark; 2003] Graeme Clark, Cochlear Implants, Fundamentals and    Applications, AIP Press, Springer Science+Business Media, Inc., New    York, N Y, 2003.-   [Bal & Oertel; 2001] Ramazan Bal and Donata Oertel, Potassium    Currents in Octopus Cells of the Mammalian Cochlear Nucleus, Journal    of Neurophysiology, Vol. 86, pp. 2299-2311, Published 1 Nov. 2001

1. A hearing assistance device comprising an implantable part forelectrically stimulating an auditory nerve of a user, the implanted partcomprising a current source generator; an electrode array configured tobe located inside one of the cochlear scala or adjacent to the auditorynerve, or at the auditory brainstem; the hearing assistance device beingconfigured to produce a time-varying waveform delivered by said currentsource generator, said time-varying waveform comprising a positivelysloping positive pulse.
 2. A hearing assistance device according toclaim 1 wherein said time-varying waveform comprising a positivelysloping positive pulse comprises a rising edge and a falling edge,wherein the height of the falling edge is larger than the height of therising edge.
 3. A hearing assistance device according to claim 1 whereinsaid time-varying waveform comprises a negatively sloping negativepulse.
 4. A hearing assistance device according to claim 1 wherein thehearing assistance device is configured to dynamically adapt thetime-varying waveform to the current input signal.
 5. A hearingassistance device according to claim 1 wherein the hearing assistancedevice is configured to provide that the time-varying waveformstimulation pulse is modulated in width and/or amplitude according tothe frequency content of a current input signal.
 6. A hearing assistancedevice according to claim 1 wherein the time-varying waveform comprisesa bi-phasic sloping, symmetric waveform stimulation pulse.
 7. A methodof operating a hearing assistance device, the hearing assistance devicecomprising an implantable part, the method comprising providing anelectrode array comprising one or more stimulation electrodes configuredto be located inside one of the cochlear scala or adjacent to theauditory nerve, or at the auditory brainstem; providing stimulationcurrent to generate electric stimulation pulses to one or more of saidstimulation electrodes; using said stimulation current to provide aparameterized time-varying waveform of said electric stimulation pulsesto one or more of said stimulation electrodes, said parameterizedtime-varying waveform comprising a positively sloping positive pulse. 8.A method according to claim 7 comprising providing a model of ioniccurrents present in a nerve or neuron, from which the temporal patternof current to deliver can be computed so that a specific dischargeprobability and/or temporal accuracy can be obtained.
 9. A methodaccording to claim 7 comprising one or more of the following steps:Computing a first passage time probability density using a model of saidparameterized time-varying waveform. Computing a set of fibres, whichare activated by a single pulse of said parameterized time-varyingwaveform. Computing interactions between sub-sequent pulses in a pulsetrain of said parameterized time-varying waveforms.
 10. A methodaccording to claim 7 comprising modulating the probability ofdischarging neurons of the cochlear nerve using a parameterizedtime-varying waveform specifically designed to limit the spread ofexcitation.
 11. A method according to claim 7 comprising modulating theparameterized time-varying waveform by acting on either the spread ofexcitation or the discharge probability or the discharge latency.
 12. Amethod according to claim 7 comprising a fitting procedure wherein thepatient estimates the extent of the spread of excitation a specificelectric stimulation pulse.
 13. A method according to claim 7 comprisinga fitting procedure, wherein a subjective measure related to the use ofa masking paradigm is provided in which the patient is asked to detectthe presence of a target stimulation in concurrence with a maskerpresented simultaneously or earlier.
 14. A method according to claim 7comprising a fitting procedure, wherein an objective measure for fittingthe pulse-shape is provided, said objective measure being based onrecording the nerve response after its stimulation or any evoked neuralresponse produced by the stimulation.
 15. A data processing systemcomprising a processor and program code means for causing the processorto perform the steps of the method of claim
 7. 16. A fitting systemconfigured to estimate the extent of the spread of excitation ofdifferent parameterized time-varying waveforms according to the methodof operating a hearing assistance device defined in claim
 12. 17. Amethod of determining a temporal pattern of a stimulation waveformcomprising: providing a model of ionic currents present in a nerve orneuron, from which the temporal pattern of current to deliver can becomputed so that a specific discharge probability and/or temporalaccuracy can be obtained; Computing the first passage time probabilitydensity using a parameterized pulse shape model; Computing the set offibres, which are activated by a single pulse of parameterized temporalshape; and Computing the interactions between sub-sequent pulses in apulse train.